Cancer, for example, oral and cervical cancer, is a growing global health problem that disproportionately impacts the developing world. Each year, over 481,000 new cases of oral cancer are diagnosed worldwide, with a 5-year mortality of ˜50% and nearly two-thirds of which occur in developing countries. Cervical cancer is the second most common cancer in women with an incidence and death rate of 16 and 9 per 100,000 women, respectively, and 80% of cases occur in the developing world. Detecting and grading precancerous and malignant oral lesions is mostly accomplished by visual screening and biopsy of suspicious tissue sites. The Pap smear is the standard of care for screening for cervical cancer. An effective cancer screening and diagnostic program often requires both sophisticated and expensive medical facilities and well-trained and experienced doctors and nurses. The high death rate in developing countries is largely due to the fact that these countries do not have the appropriate medical infrastructure and resources to support the organized screening and diagnostic programs that are available in the United States or other developed countries. Thus, there is a critical global need for a portable, easy-to-use, low cost, and reliable device that can rapidly screen for oral and cervical cancer in low-resource settings.
It is well documented in numerous studies that oral and cervical cancers, if detected at early stages, have a better chance of being successfully treated with surgery, radiation, chemotherapy, or a combination of the three, therefore significantly improving the survival rates. One such early detection method can include analyzing optical absorption and scattering properties of epithelial tissues which reflect their underlying physiological and morphological properties. UV-visible diffuse reflectance spectroscopy (UV-VIS DRS), which measures tissue absorption and scattering properties, has shown promise for diagnosis of early precancerous changes in the cervix and oral cavity. Tissue absorption and scattering can be quantified using in vivo DRS measurements. For example, in the UV and visible light band, dominant absorbers in oral and cervical tissue are oxygenated and deoxygenated hemoglobin (Hb), arising from blood vessels in the stroma. Light scattering is primarily caused by cell nuclei and organelles in the epithelium and stroma, as well as collagen fibers and cross-links in stroma. Neoplastic and cancerous tissue exhibit significant changes in their physiological and morphological characteristics that can be quantified optically: Stromal layer absorption is expected to increase with angiogenesis, whereas stromal scattering is expected to go down with neoplastic progression as extracellular collagen networks degrade. Epithelial scattering has been shown to increase, for example, due to increased nuclear size, increased DNA content, and hyperchromasia. UV-VIS DRS has a penetration depth that can be tuned to be comparable to the thickness of the epithelial layer or deeper to probe both the epithelial and stromal layers [1], [2], [3].
Hardware employed for UV-VIS DRS measurements typically consists of a broadband light source, a spectrometer for multispectral detection, and a fiber-optic probe for relaying light to and from the instrument. Although fiber-optic probes are well-suited to access tissue sites in the oral cavity and cervix they are susceptible to several sources of systematic and random error that can influence the robustness of this technology, particularly in resource-poor settings. One such error arises from an uncontrolled probe-to-tissue interface which makes it difficult to obtain a reproducible tissue reflectance spectrum due to probe to tissue coupling and physiological changes induced by the probe pressure. That is, a probe technician or operator can unknowingly interfere with spectral measurement of the specimen when contact pressure between the probe and the tissue rises. One study found that there was a decrease in the diffuse reflectance and increase in the scattering coefficient between 400-1800 nm with compression of in vitro human skin [4]. Another study reported that extracted blood vessel radius, oxygen saturation, and Mie theory slope decreased with contact pressure, while the reduced scattering coefficient at 700 nm increased as a function of pressure [5]. A more recent study concluded that elevation in probe pressure can induce major alterations in the profile of the reflectance spectra between 400-650 nm and the changes in the extracted tissue optical properties depend not only on the probe pressure, but also on tissue type [6]. It is generally believed that the changes may be attributed to the compression of the blood vessels which causes reduced blood flow and alterations in the metabolism of the tissue as well as a change in the density of the scatterers. Thus, it appears that unknown and uncontrolled contact pressure at the probe specimen interface can adversely affect measurements and early detection of affected tissue.
Another error in conventional systems can arise from the lack of a robust, real-time calibration makes the calibration process time-consuming and potentially inaccurate, particularly when attempting to quantify absolute absorption and scattering coefficients. It has been established that in order to consistently yield accurate estimation of tissue optical properties, calibration can compensate for the wavelength-dependent instrument response, lamp intensity fluctuations, and fiber bending losses [7], [8]. Conventional calibration techniques typically rely on measurements using reflectance standards of known optical properties and/or tissue phantoms, typically after the clinical measurements are completed. These measurements are subject to a number of limitations, however. First, because the calibration is performed at the beginning or end of the study, real-time instrument fluctuations, such as lamp drift and fiber bending loss cannot be compensated for. Second, these measurements can require an additional thirty minutes for lamp warm-up and another ten to twenty minutes for calibration, which adds up to a significant amount of time, especially in a clinical setting. Thus, a better calibration method is needed, on that can compensate for real-time instrument fluctuations.
Finally, and in addition to being problematic and error prone, typical DRS systems can be expensive to use as they comprise bulky, high power and expensive optical components, such as thermal light sources, spectrometers, and cooled CCD cameras, which need a stable power supply. Thermal light sources have large footprint, short life-time, low power efficiency, and low coupling efficiency to optical fibers. Spectrometers using grating spectrometers and cooled CCD cameras have extremely high wavelength resolution and sensitivity, but are very bulky and expensive and consume a large amount of electrical power. In addition, a stable power supply is very often required to operate a thermal lamp and a CCD camera. Taken together, it is very difficult for DRS systems in their current forms to be directly used for cancer screening in developing countries.
Consequently, there remains a need for improved smart fiber optic sensor systems and methods for quantitative tissue optical spectroscopy that overcome or alleviate shortcomings of the prior art systems and methods. In particular, there remains a need for a low power consumption, low-cost DRS device that can be used to obtain accurate and reproducible quantitative measurements of absorption and scattering coefficients with applications to global health screening of cervical and oral cancers. Such improvements can comprise, but are not limited to, utilizing of emitting diodes (LEDs) as illumination sources; using miniature fiber-optic spectrometers for light detection and a smart fiber-optic probe for reliable measurements of tissue diffuse reflectance spectra. The LEDs and spectrometers can be powered and controlled by a laptop computer using custom computer readable media, making the system highly portable. Smart fiber optic sensors, systems, and methods can integrate a specimen sensing channel, a self-calibration channel, and an interferometric fiber-optic pressure sensor into a single instrument. The pressure sensor can provide real-time pressure readings at the probe-specimen interface such that an operator can adjust the applied force on the probe. The spectra can only be saved and processed if the desired pressure is reached. The pressure sensor can ensure that the probe-tissue coupling is reliable and the pressure induced tissue physiological changes are consistent between measurements. The self-calibration channel can collect a calibration spectrum concurrently with the tissue spectrum, which will be used to correct for source fluctuations and fiber bending loss that occurs during the measurements. With the smart fiber optic sensors, systems, and methods disclosed herein it can be possible to perform accurate and reproducible DRS for rapid screening of cancers or charactering in vivo tissues in resource-limited countries without having to use expensive optical components and high capacity stable power supplies. More importantly, it can eliminate the need for instrument warm-up and extra on-site calibration measurements, thus saving 40-60 minutes of time.